Procedure, apparatus and detector for the determination of fractional oxygen saturation

ABSTRACT

The invention relates to a procedure for determining the relative concentration or composition of different kinds of haemoglobin, such as oxyhaemoglobin, deoxyhaemoglobin and dyshaemoglobins, and/or dye components contained in blood in a non-invasive manner using the light absorption caused by different haemoglobin varieties and/or dye components, in which procedure light signals are transmitted at at least two predetermined wavelengths to a tissue comprised in the patient&#39;s blood circulation, the light signal transmitted through the target under measurement and/or reflected from it is received and the proportion of the intensity of the pulsating light signal received at each wavelength is determined in relation to the total intensity of the light transmitted through the tissue or reflected from the tissue. In the procedure, the effective extinction coefficients of blood haemoglobin derivatives and/or dye components in the tissue are determined for each light signal and/or light signal pair via a mathematical transformation from blood dye component extinction coefficients consistent with the Lambert-Beer theory and the proportion of specific blood haemoglobin derivatives and/or dye components in relation to the total amount of haemoglobin contained in the blood is determined by means of the intensity of the signals received in different wavelength ranges.

BACKGROUND OF THE INVENTION

The present invention relates to a procedure as defined and to ameasuring apparatus for non-invasive determination of fractional oxygensaturation in blood. Moreover, the invention relates to a sensor,designed for use in conjunction with the measuring apparatus of theinvention to collect measurement data about the patient.

Specifically, the present invention relates to the monitoring of theoxygenation level of the body in patient monitoring systems. Measuringthe oxygen saturation of arterial blood in peripheral circulation isgenerally sufficient to determine the oxygenation situation andsufficiency of oxygen supply in the entire body. The oxygenation levelof the human body can be estimated via oxygen saturation measurement ofarterial blood either in a non-invasive manner using pulse oximeters ortranscutaneous oximeters/blood gas analysers or in an invasive mannereither by taking a sample of arterial blood and analysing in vitro bloodgases (In Vitro Blood Gas/pH Analyzers) or performing an opticmeasurement on the blood sample using so-called CO-oximeters orhaemoximeters (In Vitro Multiwavelength Oximeters).

Partial pressure measurements of gas in arterial blood samples and opticmethods based on the absorption of light by blood samples are part oflong-standing tradition, but clinical use of pulse oximeters only becamecommon in late 1980's and the measuring principle itself is relativelynew. There are numerous patents and patent applications relating topulse oximeters. The most important of these as well as the mostcomprehensive general descriptions of prior art are found in patentspecifications U.S. Pat. No. 4,653,498, U.S. Pat. No. 4,819,752, U.S.Pat. No. 4,407,290 and U.S. Pat. No. 4,832,484.

The prior-art technology described in the above-mentioned patentspecifications, which is the basis of currently used equipment, isimperfect and inadequate for continuous and non-invasive monitoring ofchanges in the actual oxygenation level or degree of fractional oxygensaturation in a patient's blood. Although in vitro oximeters are inprinciple capable of measuring fractional oxygen saturation from anormal blood sample, the measurement is neither non-invasive norcontinuous. On the other hand, pulse oximeters measure continuously andnon-invasively, but they are not able to measure the actual degree offractional oxygen saturation of blood and are therefore inadequate forsituations where only a part of the total amount of haemoglobin in apatient is functional. Pulse oximeters measure fractional oxygensaturation assuming that the patient's blood composition is the same asthat of a healthy, non-smoking person. A high dyshaemoglobin level, i.e.a high relative amount of haemoglobin not participating in oxygentransport, always involves a danger to the patient because current pulseoximeters produce an incorrect estimate of the oxygenation level ofblood.

The cause of incorrect measurement lies in the measuring principle:Since pulse oximeters use only two different wavelengths of light forthe estimation of oxygen saturation, only two different kinds of bloodhaemoglobin, viz. oxyhaemoglobin (HbO2) and deoxyhaemoglobin (Hb), canbe accurately measured by this method. All other dyeing blood components(usually dyshaemoglobins or dyes used in clinical tests) have adisturbing effect on the measurement and can only be taken into accountas average amounts. This type of average correction is generally made onthe composition of healthy blood. However, the composition of normalblood may change in an unforeseen manner and without a readilyidentifiable cause. The blood composition of a patient with a criticalillness may differ from the blood composition of a healthy person as aresult of medication, the nature of the illness or a medical treatmentor measurement. A new and significant treatment of this type is theso-called nitrogen oxide (NO) treatment, which may cause a considerablerise in the patient's methaemoglobin (MetHb) level. Another common caseof incorrect measurement is carbon monoxide poisoning, which involves ahigh carboxyhaemoglobin (HbCO) level in the patient. Continuousnon-invasive monitoring of the actual degree of oxygen saturation isparticularly important during NO treatment because the dyshaemoglobinlevels may rise relatively rapidly, which means that an analysis basedon a blood sample is not sufficient. The measurement of fractionaloxygen saturation is also of great importance in rescue operations andin follow-up monitoring after carboxyhaemoglobin poisoning.

It is obvious that accurate, continuous and non-invasive measurement offractional oxygen saturation requires a sensor with several wavelengthsused to produce an analysis of blood composition. In the following,prior art will be discussed by considering a technique that uses theprinciple of non-invasive measurement using an oximeter with severalwavelengths.

A previously known oximeter based on non-invasive measurement uses eightdifferent wavelengths to determine the average degree of oxygenation ofthe blood via a measurement on the ear (see Girsh et Girsh, Ann.Allergy, 42, pages 14-18, 1979). The measurement does not use the pulseoximeter principle, whereby the measurement is only applied to arterialblood by distinguishing from the light transmission a componentpulsating in synchronism with the heartbeat and normalising thiscomponent against the total light transmission. Instead, average oxygensaturation is measured directly from the total light transmission at thewavelengths used. The total transmission depends on the oxygensaturation and composition of both arterial and venous blood, but alsoon the absorption and scattering caused by other tissues. Typically,blood accounts for only 1-2% of the amount of tissue, so the signal maybe very ambiguous. Such a method has many drawbacks: First, the person'scomplexion, the structure of the tissue in itself and especially thescattering and absorption properties of the tissue as well as its otherproperties change and even dominate the total transmission. In fact, themethod requires several wavelengths for the compensation of theseproperties, and it cannot produce reliable analyses of the compositionof arterial blood. In addition, analysing the blood composition in termsof percentages is difficult because the relative amounts of arterialblood and venous blood and their different degrees of oxygen saturationaffect the absorption. The amount of dyshaemoglobins is the same in botharteries and veins, but oxygen saturation varies with tissue metabolismand temperature or with the regulating mechanisms of the body.

In patents EP 335357 and U.S. Pat. No. 5,421,329 it is suggested that byadding a third wavelength to a conventional pulse oximeter with twowavelengths it is possible to improve the accuracy of functionalsaturation measurement with a pulse oximeter or to eliminate or reducethe artifacts caused e.g. by movement. In the former patent, the thirdwavelength is used to eliminate the irregular artifacts signal from ontop of the pulsation caused by the heartbeat. The wavelength is not usedfor the determination or identification of the dyshaemoglobin level. Inthe latter patent, the third wavelength is used to adjust themeasurement of a low degree of oxygen saturation, but it is not used inconjunction with the measurement of the normal saturation range or forthe measurement of dyshaemoglobin levels or the degree of fractionaloxygen saturation. The latter patent also relates to the reflectionprinciple and especially to the measurement of oxygen saturation in ababy during childbirth. In this situation, a third wavelength isnaturally needed to achieve a more reliable measurement. Similarly,patent application WO 94/03102 proposes the use of a third wavelength toeliminate artifacts caused by motion. Patent specification U.S. Pat. No.4,714,341 (Minolta Camera) also uses a third wavelength for moreaccurate measurement of functional saturation. Like the others, thisspecification is not concerned with the measurement of fractionalsaturation or in general dyshaemoglobin levels.

Patent specification EP 0 524 083 also proposes the use of a thirdwavelength for simultaneous measurement of carboxyhaemoglobin level andoxygen saturation. In the measurement, three different laser diodes withwavelengths of 660 nm, 750 nm and 960 nm are used. These threewavelengths are used to measure the modulation ratios, and theconcentrations of three unknown kinds of haemoglobin, HbO2, Hb and HbCO,are calculated by solving a linear system of equations. However, theprocedure presented in patent specification EP 0 524 083 has twosignificant drawbacks. First, the procedure is not applicable for themeasurement of MetHb; in other words, fractional saturation can only bedetermined for the three kinds of Hb mentioned above. Secondly, solvingthe aforesaid linear system of equations is not sufficient for thedetermination of the concentrations of the aforesaid three kinds of Hb,as will become evident later on from the description of a preferredembodiment of the calculating procedure of the present invention. Thebasic drawback is that the system of equations is not a linear onebecause the coefficients used in it are in themselves functions of theconcentrations. For this reason, the use of this method is restricted toa very narrow range of oxygen saturation, and the procedure is notworkable in the operating range generally required for pulse oximeters.In addition to the above drawbacks, the procedure involves the use oflaser diodes and a fibre optic connection to the measurement point,which makes the measuring system rather too expensive for practicalmeasurements and difficult for the user. Moreover, laser diodes have anarrow choice of wavelengths. Due to the use of fibre optics, the lightis attenuated especially at the connection points and thesignal-to-noise ratio is worse than in the conventional solutionemploying light-emitting diodes. The official regulations relating tocoherent radiation and the danger caused by the radiation e.g. to theeye also constitute a limitation of the application of the procedure inpractical situations.

Further, patent specification Aoyagi et al, EP 0 679 890 A1,representing prior art, presents an apparatus designed for themeasurement of light absorbing blood components. According to thespecification, the procedure and apparatus can be used to determine thedegree of functional oxygen saturation of blood, the concentrations ofdifferent kinds of haemoglobin as well as other dye components of blood,such as bilirubin and in-vein dyes. The proposed procedure and apparatusare based on a rather unusual optical model of light transmissionthrough tissue and formation of a pulsating signal. Since the procedureis obviously one of the prior-art solutions related to the presentinvention, it is necessary to point out the erroneous assumptions lyingbehind the procedure and apparatus. The drawbacks listed below serve asexamples, and the drawbacks are not described in full extent. For a moredetailed explanation of the drawbacks, reference is made to the thesisReindert Graaff, "Tissue Optics Applied to Reflectance Pulse Oximetry",Groningen University, Feb. 12, 1993, which is an excellent descriptionof tissue optics and its modern representation. To those familiar withpulse oximetry or non-invasive measurement of blood properties, thedrawbacks listed below are self-evident and can be recognised viaempirical measurements. Accordingly, patent application EP 0 679 890 isbased on the following, erroneous propositions. First, diffusionapproximation and its parametrised flux models (in the applicationreferred to, the so-called Arthur Schuster theory) can be applied todescribe the total transmission through tissue, but they cannot be usedin conjunction with pulsating tissue components and the operation ofpulse oximeters at short wavelengths (600-700 nm), nor can theygenerally be used at a low saturation or for highly absorptive blood dyecomponents. Using the diffusion model (applies to situations where thescattering cross-section is considerably larger than the absorptioncross-section) together with the Lambert-Beer pulse oximeter model(applies to situations where the absorption cross-section isconsiderably larger than the scattering cross-section) simultaneouslygenerally does not lead to realistic results. Second, it is stated inthe patent application that the scattering term is known and independentof the wavelength. In fact, the scattering term is one of the adjustableparameters in the model and is also dependent on the wavelength and thetissue type. The pulsating portion of the scattering is also dependenton the size, shape and number of blood cells, i.e. on the haematocrit.Third, the tissue term (in the patent application, the pulsatingcomponent that is not blood) plays no significant role in the signalformation at all and therefore it cannot be used in the way it appearsin the formulas as a factor representing theoretical extinctioncoefficients of blood and empirically measured modulation ratios.Further, it is stated in the patent application that the tissue term ismostly water, which in fact does not absorb at all in relation to thedominating pulsating terms in the wavelength range used. In fact, thedominating pulsating tissue-type effect is produced by blood, whoseabsorption depends on all those things that patent application EP 0 679890 presents as quantities to be measured, such as oxygen saturation,different haemoglobin varieties and their amounts and dyes; thus, therewould be no linear correlation between different tissue terms that couldbe defined in advance--although the application asserts to thecontrary--which means that the degree of non-linearity of the problemincreases considerably. Fifth, the number of unknown quantities in theprocedure and apparatus of the patent application in question clearlyexceeds the number of equations available. Moreover, the patentapplication contains numerous other inaccuracies, so the application orthe procedure and apparatus presented in it do not, at least in respectof their basic assumptions, meet the quality criteria that are expectedto be observed in clinical patient monitoring measurements.

In the above, existing prior art has been dealt with from the point ofview of systems using more than two different light sources each havinga different spectral emission, yet so that the spectral emission in thesame light source is always the same. Especially the use of more thantwo light sources is still associated with problems relating tomaintaining the accuracy of the apparatus in use even in situationswhere the spectral emission of the light sources changes due totechnical aspects of fabrication of the light source and maintainingaccuracy requires a correction to compensate this change. A method forsuch correction or rather a cheap method for maintaining sensor accuracyis presented in U.S. Pat. No. 4,621,643 (December 1986), U.S. Pat. No.4,700,708 (October 1987) and U.S. Pat. No. 4,770,179 (September 1988).All these patents propose solutions in which information about thewavelengths of the light sources is transmitted to the measuringapparatus by encoding the correction required by changes of wavelengthinto an impedance element or in practice into the resistance value of aresistor. From the resistance value or by some other similar codingmethod, the measuring apparatus receives information indicating thechanges required in the calibration of the apparatus. This can be donewith a single resistance value or other simple `coding` when theunambiguity of the measurement signal is guaranteed via othertechniques. In two-wavelength pulse oximeters, unambiguity is based onthe apparatus forming substantially only one signal, i.e. a modulationratio between the two wavelengths, a so-called R-value, which, via anunambiguous calibration curve, can be directly associated with thefunctional oxygen saturation or SpO2 value. No such unambiguouscorrelation exists when there are several light sources and more thantwo haemoglobin varieties or other blood dye components are to bemeasured. As a summary of prior art, it can be stated that so far thereis no method or apparatus capable of reliable measurement of fractionaloxygen saturation of arterial blood and quantitative determination ofthe dyshaemoglobin level.

SUMMARY OF THE INVENTION

The present invention proposes a different calibration method, which isunambiguous and is not--unlike prior-art technology--based ontransmitting the emission properties of the light sources to a measuringapparatus but is instead based on light source-specific use of theabsorption properties of each haemoglobin component or dye, i.e. theso-called extinction coefficients of blood. This allows e.g. situationswhere the wavelengths of the light sources or the light source typeitself can be changed or the sensor can be aligned for different kindsof Hb or dye measurements while preserving the compatibility andaccuracy of the sensor with all apparatus. Thus, in the framework of thepresent invention, the "wavelength" of the light source is understood inthe first place as meaning the current number or other identification ofthe light source, which is associated with the haemoglobin components ordyes measured with the sensor in question. The new method is applicableboth for sensors and apparatus with two light sources and for those withmore than two light sources.

The object of the present invention is to eliminate the problems andinaccuracies described above. A specific object of the present inventionis to produce an effective and accurate measuring procedure for thedetermination of the relative concentrations or compositions ofhaemoglobin derivatives or dye components contained in a patient'sblood. A further object of the present invention is to produce ameasuring apparatus and a sensor which can be utilised to effectivelyapply the calculation method of the present invention for determiningthe level of oxygen saturation in a patient's blood.

As for the features characteristic of the present invention, referenceis made to the claims.

In the procedure of the invention for determining the relativeconcentration or composition of different kinds of haemoglobin containedin blood, such as oxyhaemoglobin, deoxyhaemoglobin and dyshaemoglobins,and/or dye components, such as various in-vein dyes or the like, in anon-invasive manner using the light absorption caused by differenthaemoglobin varieties and/or dye components, light signals aretransmitted at at least two predetermined wavelengths to a tissuecomprised in the patient's blood circulation, a light signal transmittedthrough the target under measurement and/or reflected from it isreceived and the proportion of the intensity of the pulsating lightsignal received at each wavelength is determined in relation to thetotal intensity of the light transmitted through the tissue or reflectedfrom the tissue. The pulsation of the light signal is determined by theheartbeat frequency, which has a direct effect on the amount of bloodflowing in the tissue and therefore also on the amount of haemoglobinderivatives and/or dye components.

According to the invention, the effective extinction coefficients ofblood haemoglobin derivatives and/or dye components in the tissue aredetermined for each light signal and/or light signal pair via amathematical transformation from the extinction coefficients of theblood dye components according to the Lambert-Beer theory and theproportion of specific blood haemoglobin derivatives and/or dyecomponents in relation to the total amount of haemoglobin contained inthe blood is determined by means of the intensity of the signalsreceived at different wavelengths. Thus, the procedure of the inventionis based on a so-called modulation signal for each wavelength and on acomparison of these signals between two different wavelengths. Theresult of the latter comparison is expressed in terms of a modulationratio. This quantity describes the average blood dye difference betweenthese two wavelengths. When this relative dye difference is measuredusing several wavelength pairs, the concentrations of differenthaemoglobin derivatives are obtained by solving a non-linear system ofequations (1) formed from the wavelength pairs and extinctioncoefficients. ##EQU1## where % mod i is the modulation percentage forlight transmission as measured at wavelength i, i.e. the proportion oflight transmission varying at heartbeat frequency as a percentage of thetotal light transmission;

C is a constant;

the ij-element of the T (ε)-matrix is an empirical extinctioncoefficient of blood haemoglobin variety and/or dye component HbX_(j),mathematically derived from the known extinction coefficient forwavelength i; and

the unknown kinds of blood haemoglobin and/or dye components inpercentages are placed in the vertical vector (HbX₁, HbX₂, . . . ,HbX_(j)) . The non-linearity of the system of equations is due todivergences between theory, and practice. The actual extinctioncoefficients ε' are also dependent on the scattering of light caused bythe tissue and on the combined effect of absorption and scattering. Thecorrections needed in the extinction coefficients are larger the largeris the proportion of the attenuation caused by absorption andscattering. In the Lambert-Beer theory, the scattering and the effect ofthe tissue are not taken into account.

The system of equations presented above can be solved in severaladvantageous ways according to the present invention. In an embodiment,the system of equations is solved for all blood dye components, the sumof whose proportions is 100%, and the number of independent lightsignals is selected so that it corresponds at least to the total numberof haemoglobin derivatives and/or dye components set as unknowns.Moreover, it is to be noted that blood may also contain haemoglobinderivatives or dye components whose concentration and/or composition isknown. In this case, the known concentrations must be taken into accountwhen calculating the total number of all haemoglobin varieties and/ordye components.

In an embodiment of the present invention, the haemoglobin varieties setas unknowns are oxyhaemoglobin and deoxyhaemoglobin, and at least onedye component is a blood dyshaemoglobin variety, such as HbCO, MetHb orHbNO.

In an embodiment of the present invention, the non-linear system ofequations is solved by using modulation ratios, wherein thetransformation between the known extinction coefficients according tothe Lambert-Beer theory and the effective extinction coefficient of theblood-containing tissue is a function transformation between themeasured modulation ratio of two independent light signals and themodulation ratio formed from corresponding known extinctioncoefficients.

In an embodiment of the present invention, the non-linear system ofequations is solved by dividing the non-linear system of equations intolinear portions on the basis of blood composition and around a givencomposition and solving the linear system of equations thus obtainedusing experimentally determined extinction coefficients derived for theblood composition in question. The extinction coefficients arepreferably determined in advance for different blood oxygenation levels.The non-linear system of equations thus reverts into a linear system ofequations and it can be easily solved by known mathematical methods.

In a preferred embodiment, the non-linear system of equations is solvedvia an iterative process, part of which process consists in identifyingthe composition and/or existence of the dyshaemoglobin and/or dye thatleads to the best iterative result. Iteration is a known mathematicalmethod and is therefore not explained here in detail.

Further, in the determination of the concentrations of haemoglobinvarieties and/or dye components, it is preferable to weight differentmodulation ratios on the basis of the patient's blood composition, inwhich case the non-linear system of equations (1) is solved iterativelyby using weighted modulation ratios.

In a preferred embodiment of the present invention, a light signal istransmitted via the same or nearly the same optic route to the tissueusing at least two predetermined wavelengths. As the light signal passesthrough the tissue at substantially the same point, errors due todivergences in the tissue are avoided.

The selection of wavelength ranges is explained below in greater detailby referring to FIG. 1, but in a preferred embodiment the wavelengths tobe used are selected from four different wavelength ranges, of which atleast one wavelength range is below 660 nm, preferably so that thewavelength ranges are: 620-650 nm, 655-665 nm, 680-750 nm and 790-1000nm. The selection of wavelength ranges is further affected by thecomposition of human blood so that for a normal blood composition wherethe total amount of dyshaemoglobin is below a suitable level, e.g. below3%, light signals produced at center wavelengths of 660 nm, 690 nm and900 nm and/or modulation ratios formed from them are weighted inrelation to the light signal in the wavelength range of 620-650 nm andits modulation ratios. For the measurement of a high methaemoglobinlevel, the wavelength range of 620-650 nm and its modulation ratios areused, which are weighted in substantially the same proportion as theother light signals used and their modulation ratios. In addition, for anormal composition of human blood, the measurement can be performed atfour center wavelengths 900±10 nm, 690±5 nm, 658±5 nm and 632±5 or900±10 nm, 690±5 nm, 658±5 nm and 645±5 nm.

In a preferred embodiment of the present invention, for each independentlight signal an effective extinction coefficient is calculated in thelight signal emission range concerned from wavelength correlations ofknown extinction curves for blood haemoglobin varieties and/or dyecomponents. Further, an effective extinction coefficient in the tissueis determined via the same mathematical transformation for all smallspectral changes in a given light signal.

In a preferred embodiment of the present invention, changes in thespectral emission of a light source operating around a given centerwavelength, i.e. variations in the wavelength of the light source aretaken into account by utilising the changes in the effective extinctioncoefficient consistent with the Lambert-Beer theory, by determining themodulation ratio between two light signals for the effective modulationratio measured in the tissue and the effective modulation ratioaccording to the Lambert-Beer theory and determining the concentrationsof haemoglobin varieties with the aid of the corrected effectiveextinction coefficients and the effective modulation ratio.

In a preferred embodiment of the present invention, the measurement iscalibrated separately for each sensor by storing the determinedeffective extinction coefficients consistent with the Lambert-Beertheory and/or corrected to correspond to the tissue separately for eachlight source, preferably for each light element, in a storage deviceprovided in the sensor so that they can be used in the determination ofhaemoglobin varieties and/or dye components. Furthermore, the functiontransformations for each light signal pair and/or the center wavelengthsare stored in a storage device in the sensor, to be used in thedetermination of haemoglobin varieties and/or dye components.

In a preferred embodiment of the present invention, to determine a givendyshaemoglobin variety, the proportion of other dyshaemoglobin varietiesis set as a constant and the variety to be measured is set as anunknown. In addition, the proportion of other varieties is set to avalue corresponding to an invasively measured proportion.

The invention also relates to a sensor for the collection of measurementdata through tissue comprised in a patient's blood circulation innon-invasive measurement, said sensor comprising means for connectingthe sensor to a measuring apparatus, a light source which transmits alight signal at at least two predetermined center wavelengths and areceiver disposed to receive a light signal transmitted through and/orreflected from the target under measurement. According to the invention,the sensor comprises a storage device for the storage of predeterminedsensor-specific data, where the sensor-specific data comprises aneffective extinction coefficient for each haemoglobin variety and/or dyecomponent desired, said effective extinction coefficient beingcharacteristic of each light source used. The storage device may be aprogrammable read-only memory whose contents can be electrically deletedor altered, or a similar memory circuit.

It is further preferable that in the storage device are stored thefunction transformations and/or center wavelengths for each light signalpair and/or an identifier representing these and/or other correspondinginformation that indicates the connection between extinction accordingto the Lambert-Beer theory and extinction measured in tissue, to be usedfor the determination of haemoglobin varieties and/or dye components.

In a preferred embodiment of the sensor of the present invention, thelight source comprises a set of light elements in which the wavelengthsto be used, of which at least one is below 660 nm, have been selectedfrom four wavelength ranges as follows: 620-650 nm, 655-665 nm, 680-750nm and 790-1000 nm. Further, the light elements are thermally anchoredon the sensor frame to keep the temperature of the light elements belowa specific limit. Moreover, to keep the temperature of the sensor partto be connected to the patient below a specific limit, the sensorcomprises a first optic fibre for passing the emitted light to thetarget under measurement and a second optic fibre for passing thereceived light to the receiver. In this way, the elements, which tend toget warmed up, can be kept at a distance from the patient.

In a preferred embodiment of the present invention, the sensor comprisesa first set of light elements arranged to emit light to a first targetand a second set of light elements arranged to emit light to a secondtarget. Further, the first and second sets of light elements have onecommon wavelength range, which is used for the compensation ofvariations due to the measuring point. In a preferred case, themeasurement signals obtained in the common wavelength range are comparedwith each other and, based on this comparison, divergences caused by thetissue between measuring points are adjusted.

The sensor may also preferably comprise a set of light filters disposedin conjunction with the receiver to divide the received light intodifferent wavelength ranges. In this manner, light can be transmitted ina wide range of wavelengths, preferably in the range of 600-1000 nm, andthe received signal can be filtered into desired wavelength bands.

The sensor may further comprise fixing means for attaching the sensor tothe patient, preferably to the patient's ear or finger. It is alsopossible to attach the sensor to other limbs or organs of the patient.

The present invention also relates to a measuring apparatus for thedetermination of the relative concentrations or compositions ofdifferent kinds of haemoglobin contained in blood, such asoxyhaemoglobin, deoxyhaemoglobin and dyshaemoglobins, and/or dyecomponents, in a non-invasive manner using the light absorption causedby different kinds of haemoglobin and/or dye components. The measuringapparatus of the invention comprises a sensor as described above and asignal processing device for the processing of the signals received. Thesignal processing device may be a computer, a microprocessor or anapplication-specific integrated circuit (ASIC) or the like. According tothe invention, the apparatus comprises a calculating device and a readerdevice connected to the calculating device and sensor for reading thedata stored in the sensor and transmitting the data to the calculatingdevice. The reader and calculating device may be any electricallycontrolled component or application known in itself, and it canpreferably be incorporated in the signal processing equipment or in thesame assemblage with said equipment.

In a preferred embodiment of the present invention, the apparatus isarranged to measure at different wavelengths in accordance with apredetermined time division principle in such manner that themeasurement of the shares of certain dye components or haemoglobinvarieties is weighted with respect to time so that in a given period oftime the apparatus is only measuring in a given part of the entirewavelength range in use. The time division can be effected e.g. using asuitable channel arrangement and channelling device by selecting a givenchannel at a time for measurement.

In an embodiment of the measuring apparatus of the present invention,the data stored in the sensor includes specific wavelength values forspecial blood compositions of patients, essential spectral emissioninformation about the light sources used in the sensor, preferably lightelements, with respect to different dye components and/or haemoglobinvarieties, and/or information about the extinction coefficientsaccording to the Lambert-Beer theory and about the mathematicaltransformation between these and the extinction coefficients for thetissue. Further, the data stored in the sensor includes informationdescribing the sensor type. Moreover, the measuring apparatus maycomprise an identifier device for the identification of sensor type. Theidentifier device may also be integrated with the reader device.

As compared with prior art, the present invention has the advantage thatusing the procedure and apparatus of the invention it is possible toeliminate the problems described above relating to prior-art equipment,and above all the invention makes it possible to eliminate the drawbacksand inaccuracies of prior-art methods and apparatus.

Furthermore, the invention presents a new type of calculating methodthat takes into account the inaccuracies resulting from the scatteringand absorption caused by tissue. In addition, the procedure allowscompensation of small variations in LED wavelengths.

BRIEF DESCRIPTION OF THE SEVERAL VIEWS OF THE DRAWING

In the following, the invention is described by the aid of a fewexemplary embodiments by referring to the attached drawing, in which

FIG. 1 presents the extinction coefficients of different haemoglobinvarieties as functions of wavelength;

FIGS. 2a-2c represents a preferred wavelength selection according to thepresent invention;

FIG. 3 is a diagram representing a sensor according to the presentinvention;

FIG. 4 is a diagram representing a preferred arrangement for mounting alight element according to the present invention;

FIG. 5 is a diagram representing a sensor according to the presentinvention;

FIG. 6 is a diagram representing a measuring apparatus according to thepresent invention;

FIG. 7 illustrates the correlation of theoretic and empirical oxygensaturation as functions of the ratio (R/IR) of modulation percentagesdetermined at wavelengths 660 nm (R) and 940 nm (IR).

DETAILED DESCRIPTION OF THE INVENTION

FIG. 1 shows the extinction coefficients of different haemoglobinvarieties as functions of wavelength. Referring to FIG. 1, certainreasons lying behind the choice of wavelength according to the inventionwill now be explained. Preferred wavelengths are selected mainly usingtwo criteria: One of the wavelengths, e.g. λ₂, comprised in themodulation ratio (%modλ₁ /%modλ₂) is selected from a range near theisobestic point of the dominating Hb variety or oxyhaemoglobin (HbO2)and the Hb variety whose amount is primarily being measured with theratio in question, or in general from a range where a change in therelative amount of haemoglobin has only a minor effect on the signal.The other wavelength in the modulation ratio, in this case λ₁, is soselected that the Hb varieties in question have a large differencebetween their extinction coefficients at this wavelength. If theisobestic point cannot be used, it is preferable to select thewavelengths so that the difference between the extinction coefficientsof the two primary Hb varieties is of opposite sign at the twowavelengths used.

The selection of the primary modulation ratio used in the calculationwith four wavelengths from all the ratios possible or all six possiblepairs of two wavelengths is also made on two grounds: The modulationratio must have a high or at least a sufficient sensitivity to a changein the concentration of the Hb variety (Hbxx below) primarily to bemeasured with this modulation ratio; i.e. ##EQU2##

is large enough.

On the other hand, the sensitivity of the modulation ratio to sensorspecific wavelength variation between the LED components must be low inrelation to sensitivity, i.e. ##EQU3##

is small enough.

Even for as few as four haemoglobin varieties or dye components,simultaneous satisfaction of all criteria is impossible. Selecting thewavelength near the minimum or maximum of the absorption curve reducesthe calculation errors caused by variation in the wavelengths of theLEDs. For this reason, the minima and maxima of the absorption curvesand in general their flat portions are particularly good choices. Ifsuch a flat portion lies near the isobestic point, the wavelength inquestion is a good reference value, against which the dye differencesare formed. In FIG. 1 there are two such ranges: in the range 790-1000nm, the absorption curves for all Hb varieties are sufficientlyinsensitive to changes in wavelength, whereas in the range 680-750 nm,only the absorption curve of deoxyhaemoglobin is sensitive to wavelengthchanges. It is generally advisable to select the wavelengths so that twoof the four or more possible wavelengths are the same as in currentlyused pulse oximeters. When this is the case, all the empiricalinformation available in the case of pulse oximeters is also availablewhen fractional oxygen saturation is measured by means of modulationratios. The third wavelength range selection is therefore 655-665 nm.Selecting the wavelengths in these three ranges improves the accuracy offractional oxygen saturation measurement on normal human blood. Thewavelength ranges mentioned are also suitable for the determination ofHbCO.

As it is not possible to make selection of four wavelength rangeswithout conflicts, MetHb is treated as a special case. For MetHbmeasurement, a preferred wavelength range is 625-650 nm. The selectionof four different wavelength ranges according to the present inventionis: 625-650 nm, 655-665 nm, 680-750 nm and 790-1000 nm. The calculationprocedure is then so adjusted that in all situations a maximal accuracyof determination of fractional oxygen saturation is achieved for thesewavelength ranges.

In FIGS. 2a-2c, a preferred optimal wavelength selection is consideredusing the Lambert-Beer model. Pulse oximeter type modulation signals arepresented for three different special situations: 1. Hypoxemia or highHb level with normal DysHb levels (FIG. 2a), 2. Carboxyhaemoglobimia orhigh HbCO while other Hb concentrations are normal (FIG. 2b) and 3.Methaemoglobimia or high MetHb level with normal blood composition inother respects (FIG. 2c). In the case presented in FIG. 2, a preferredoptimal wavelength selection according to the invention is 900±10 nm,690±5 nm, 658±5 nm and 630±5 nm. Since the isobestic point betweenoxyhaemoglobin and carboxyhaemoglobin lies near 645 nm, this is apreferable wavelength for the calculation of the modulation ratio aswell. At this wavelength, MetHb can be determined with sufficientaccuracy. Another optimal wavelength selection according to theinvention is thus 900±10 nm, 690±5 nm, 658±5 nm and 645±5 nm. Table 1below further illustrates the contradictory nature of the wavelengthselection when maximal accuracy of determination is requiredsimultaneously for all different situations. Table 1 shows thosewavelengths which should be used in the first place, as well as thosewhich should not be used, for each one of the different descriptions ofthe patient's condition regarding oxygenation.

                  TABLE 1                                                         ______________________________________                                                                         Wavelengths                                            Wavelengths  Wavelengths                                                                             that should                                  Special   that should  that can be                                                                             not be used                                  case      be used (nm) used (nm) (nm)                                         ______________________________________                                        Hypoxemia 900, 690, 658                                                                              645       630                                          Carboxyhaemo-                                                                           900, 690     658, 645  630                                          globimia                                                                      Methaemo- 900, 690, 630                                                                              645       658                                          globimia                                                                      ______________________________________                                    

According to the present invention, the contradictions in the selectionof wavelength can be reduced and a better selection accuracy can beachieved by weighting the correct modulation ratios in different ways ineach blood oxygenation situation. For instance, in the case ofhypoxemia, wavelengths 900 nm, 690 nm and 658 nm and modulation ratioscalculated from these are always used and the fractional oxygensaturation is weighted more than the quantities for wavelengths 630 nmor 645 nm. In this case, wavelength 630/645 nm is primarily used todetect the presence of MetHb, but the measurement of the amount of DysHbis effected using other wavelengths. A preferable wavelength pair is 690nm and 900 nm because both HbCO and MetHb affect this modulation ratioin the same way. When the level of MetHb is found to be rising, theweighting coefficients are altered so that the weighting of wavelength658 nm is reduced and that of wavelength 630/645 nm is increased. In thecase of carboxyhaemoglobin, all wavelengths can be weighted moreequally. However, since the absorption coefficients forcarboxyhaemoglobin are exceptionally small in the whole range, the HbCOconcentration is mainly evident through the fact that carboxyhaemoglobinreplaces oxyhaemoglobin. The best wavelength to allow this to bedetected is 900 nm, so the modulation ratios for this wavelength must beweighted above the average. A preferred calculation method according tothe invention is weighted calculation as described below, which isoptimised for the wavelengths used and for the patient's illnesscondition. The present invention also uses a sensor calibrationprocedure to be described later on, which allows a good analysingaccuracy to be maintained at all wavelengths.

FIG. 3 presents a preferred sensor according to the present invention,which is used in the procedure of the present invention for collectingmeasurement data in a non-invasive measurement through tissue comprisedin the patient's blood circulation. The sensor comprises means 1 forconnecting the sensor to a measuring apparatus. In this embodiment, thesensor is connected to the measuring apparatus via a cable, known initself, designed for the transmission of signals. Furthermore, thesensor comprises a light source 2 forming an essential part of it, whichemits a light signal at at least two, in the disclosed embodiment, fourpredetermined medium wavelengths. The light source comprises a number oflight elements 2¹, . . . , 2⁴, each one of which emits light at aselected wavelength different from the others. The sensor also comprisesa receiver 3, which typically is a light-sensitive diode or a so-calledPIN diode and which is arranged to receive a light signal transmittedthrough and/or reflected by the target under measurement. The sensorfurther comprises a storage device 4 for the storage of predeterminedsensor-specific data, where the sensor-specific data comprises anextinction coefficient separately determined for each light source andeach blood Hb variety and/or dye component that can be measured. Thesensor presented in FIG. 3 further comprises a sensor terminal 14, inwhich the storage device is mounted and to which the light elements 2¹,. . . , 2⁴ and the receiver 3 are connected. Via the sensor terminal 14,it is possible e.g. to control the sensor by means of the measuringapparatus and to read the data contained in the storage device 4.

The light elements 2¹, . . . , 2⁴ in FIG. 3 are connected in a bipolarcircuit, which uses only three conductors instead of five and istherefore simpler and more advantageous in respect of cable structure.Light elements 2¹, . . . , 2² and, on the other hand, light elements 2³,. . . 2⁴ are connected in reverse directions relative to each other andthey are driven by operating current of opposite sign. Such a bipolarcontrol circuit in itself is obvious to the person skilled in the artand is therefore not described here in detail. Generally, the lightelements 2¹, . . . , 2⁴ are small LED chips about 0.3*0.3 mm² in size,which are adjusted on a hybrid or circuit board or a correspondingmounting. Surface mounted devices can also be used, but to produce alight source unit as small as possible, chips are a better choice.Further, one light element 2^(n) may consist of two identical lightsources (not shown) connected in series. Such a connection is preferableif the LED has an insufficient luminosity when used alone or if theattenuation caused by the tissue at this wavelength is so large that thelight transmitted through it is of insufficient intensity.

FIG. 4 presents a particularly advantageous solution that allows thesurface temperature of the sensor to be reduced in relation to theconventional structure. In this embodiment of the invention, a lightsource hybrid 2¹ is thermally anchored on the sensor frame 5 and itsthermal connection to the skin surface is minimised. Such a structureallows a low sensor surface temperature to be maintained even if severallight elements are used. Moreover, in the structure presented in FIG. 4,there is between the sensor surface and the light source hybrid a cavity15 that causes diffuse reflection, thus diffusing and smoothing thelight emitted. Therefore, the light emission reaching the skin surfaceis smoothly spread over a large area, which is an advantage e.g. inregard of the elimination of motional artifacts from the light detectorsignal. Comprised in the sphere of the present invention are alsosensors that may use different wavelength ranges and be intended forsome other measurement purpose. Such a sensor may be e.g. one intendedfor the measurement of blood bilirubin or a given in-vein dye.

FIG. 4 also illustrates a preferred embodiment of the invention, inwhich the receiver 3 is disposed on the same side relative to the tissueas the light element 2¹, so that the receiver receives a signalreflected from the tissue. The sensor depicted in FIG. 4 is particularlyadvantageous in cases where the attenuation caused by the tissue is sostrong that the light transmission at all wavelengths is insufficientfor the measurement of fractional oxygen saturation in the transmissiongeometry. In such cases, a reflection sensor as presented in FIG. 4 canbe used instead of a series connection of a number of LEDs. Since lighttransmission decreases especially in the case of short wavelengths, buton the other hand scattering increases at these wavelengths, theadvantages of the reflection geometry become manifest via bothmechanisms. Thus, the signal produced by the reflection sensor increasesin relation to other sensor solutions.

FIG. 5 presents a typical finger sensor structure, which comprises afirst optic fibre 6 for passing a light signal emitted by a light source(not shown) to the target under measurement and a second optic fibre 7for passing a light signal transmitted through the tissue to thereceiver (not shown). With this sensor structure, too, excessivetemperatures against the patient's skin are avoided. In the fingersensor, the fastening device 9 used to attach the sensor components nearthe target to be measured is a clothes-peg type clamp. Let it be furtherstated that the solution illustrated by FIG. 5 can be implemented e.g.as an ear sensor or the sensor components can be fastened by means oftape or some other type of separate fixing means.

When a fibre-optic circuit as shown in FIG. 5 is used, the LEDs as wellas the detector may be placed inside a patient monitor or acorresponding measuring apparatus, in which case the light is passed viathe first optic fibre 6 to the sensor and via the second optic fibre 7back to the measuring apparatus. In this solution presented in FIG. 5,sensor temperature rise is avoided altogether.

In a preferred embodiment (not shown) of the present invention, twoseparate sensors are used at different measuring points. One of thesensors is preferably a conventional pulse oximeter sensor. For theother sensor, one common wavelength is selected and used to correct thescaling of the modulation ratios between the sensors, calculated in acrosswise manner. Therefore, in two separate sensors, at least five LEDsmust be used instead of four. However, this structure is preferable evenin a case where one of the sensors is only used for short periods tomeasure the dyshaemoglobin level while the other sensor is usedcontinuously e.g. for the measurement of functional oxygen saturation.

Referring again to FIG. 3, the light detector 3 or receiver can bedivided into four separate wavelength channels by using light filters.FIG. 3 shows light filters 8¹, . . . , 8⁴ arranged between the tissueand the receiver 3. Each filter has its own wavelength pass band inaccordance with the wavelengths used in each measurement. In a preferredembodiment, the light filters can be electrically controlled so that thepassband wavelength can be changed by means of the measuring apparatus.When light filters are used, it is preferable to use a common wide-bandlight source. In principle, such a light source with sufficientbandwidth may consist of one or more LEDs with a center wavelength ofabout 680 nm and a bandwidth of about 60-80 nm. In this case, the LEDfor the near-infrared range may have a wavelength of e.g. 910 nm.Another solution is to use a wide-band halogen or other conventionallight source and pass the light via fibre to the oxygen saturationsensor.

FIG. 6 presents a preferred embodiment of the measuring apparatus of thepresent invention. The measuring apparatus in FIG. 6 comprises a sensoras described above by referring to FIG. 3, together with a sensorterminal 14. The measuring apparatus preferably comprises at least asignal processing device 10, which may be a microprocessor or acorresponding programmable component known in itself. Further, in thisembodiment the light elements 2¹, . . . , 2⁴ are controlled by twobipolar drive circuits (not shown) controlled by a microprocessor. Theradiation emitted by the light elements 2¹, . . . 2⁴ is passed throughthe tissue or reflected from it to the light detector 3 of the sensor,from where the signal is passed via the sensor terminal 14 to thecurrent-voltage converter (not shown) of the preamplifier of theapparatus. After this, the signal is amplified in a manner controlled bythe microprocessor 10. Let it be further stated that in FIG. 6 themeasuring apparatus and sensor are presented in a greatly reduced formbecause most of the technology used in the measuring apparatus consistsof electronics known to the person skilled in the art. The userinterface and display functions as well as other general properties ofthe measuring apparatus are defined via the microprocessor 10. Accordingto the invention, the measuring apparatus presented in FIG. 6 furthercomprises a calculating device 11, which is programmed to carry out themeasuring procedure of the invention with the aid of the sensor. Otheressential components in the measuring apparatus are a reader device 12and an identifier device 13, by means of which the information stored inthe memory element 4 of the sensor is read and the sensor typeidentified. In this context it should also be noted that with moderntechnology even a hardware implementation is possible in which all theabove-mentioned components, i.e. the signal processing device 10,calculating device 11, reader device 12 and identifier device 13, areprogrammed into a single application-specific integrated circuit (ASIC).

A preferred embodiment of the present invention is one in which themeasuring time in the embodiment presented in FIG. 6 is flexibly dividedbetween the channels. In this case, e.g. most of the time would be spenton measuring the fractional oxygen saturation while a clearly shortertime is reserved for dyshaemoglobin measurement or the dyshaemoglobinlevel is only determined when necessary. The time division used in themeasurement may be a fixed division or it can be flexibly changed asrequired in each situation.

The theory of pulse oximetry is generally presented as being based onthe Lambert-Beer law. According to the theory, light transmissionthrough the tissue at each wavelength is exponentially dependent on theabsorbance of the tissue. This theory is generally accepted andestablished in pulse oximetry. Omitting details, the theory isgeneralised for four different wavelengths in the matrix formatdescribed above, as follows: ##EQU4## where %mod i is the modulationpercentage for light transmission as measured at wavelength i, i.e. theproportion of light transmission varying at heartbeat frequency as apercentage of the total light transmission, the ij-element of theε-matrix is the extinction coefficient of the haemoglobin component j ofarterial blood for wavelength i and the haemoglobin components inpercentages are placed in the vertical vector (HbX₁, HbX₂, . . . ,HbX_(j)) in this order j. Thus, the horizontal lines of the ε-matrix arethe extinction coefficients of a given wavelength for different Hbvarieties. The constant C determines the units on the left side of theequation.

In the above system of equations it has been assumed that the modulationpercentage is small (<10%). If this is not the case, instead of %mod ithe exact theoretic form is used:

    % mod i←ln(l±% mod i/100%)*100%

The modulation percentages are the basic signals of the pulse oximeter,so in the Lambert-Beer theory it should be possible to calculate theconcentrations of different Hb varieties directly by using a matrixreverse to the ε-matrix. However, in practice this system of equationsis not linear. As stated before, the divergences between theory andpractice are due to the fact that the actual extinction coefficients ε'are also dependent on the scattering of light caused by the tissue andblood and on the combined effect of absorption and scattering. Thecorrections are larger the larger is the proportion of the attenuationcaused by absorption and scattering. The Lambert-Beer law assumes thatthe scattering of light and the non-homogeneity of tissue are not takeninto account. In practice, therefore, in non-invasive measurement theelements ε_(ij) of the ε-matrix differ from the extinction coefficientsε'_(ij) of real blood. This correlation between the actual andtheoretical extinction matrices can be represented by a transformation Tthat depends on the total absorption and scattering at the wavelengthsused in measurement. This relationship can be presented in the followingform: ##EQU5##

The * sign at the top edge of the matrix means that the elements of thematrix are ε'_(ij).

Since the total absorption is the sum of the extinction coefficientsweighted by the haemoglobin proportions, the transformation T isdependent on the haemoglobin concentrations themselves, or

    (T)=(T)(HbX.sub.1, HbX.sub.2, HbX.sub.3, . . . , HbX.sub.j)

In multi-component analysis of blood composition, the divergence betweentheory and practical measurements has significant consequences, whichare disregarded e.g. in patent application EP 0 524 083 A1. First, thesystem of equations (1) is not linear, i.e. it cannot be solved by usinga reverse matrix. Second, any Hb variety and change in its concentrationalso affects the actual extinction coefficients of other Hb varieties.For this reason, the system of equations (1) must be solved for all Hbvarieties present in blood, and the fairly strongly absorbing MetHbvariety cannot be excluded from the analysis.

In the following, referring to FIG. 7, which depicts the correlation oftheoretic and empiric oxygen saturation as functions of the ratio (R/IR)of the modulation percentages determined at wavelengths 660 nm (R) and940 nm (IR), it will be demonstrated how the system of equations (1) issolved with the aid of the modulation percentage ratios for several Hbvarieties, i.e. HbO2, Hb, HbCO, MetHb and Hbx, where Hbx is ahaemoglobin component that appears in the patient's blood in a specialsituation, e.g. nitrosylhaemoglobin HbNO or sulphohaemoglobin HbS. Inprinciple, Hbx may represent any blood dye component, such as anartificial dye.

In FIG. 7, the significance of the matrix transformation (2) is that thetransformation serves to transfer the theoretical Lambert-Beer curveonto the empirical curve (in the figure, Wukitsch et al). This can beeffected via a function conversion that changes the numeric values onthe R/IR axis in the desired (non-linear) manner. With four differentwavelengths, a corresponding correlation between the theoretical andempirical modulation ratios and the oxygen saturation level is obtainedfor each one of the six different wavelength pairs. The transformation Tcontains the information needed for the transfer of all these pairs ofcurves, and conversely, the transfer of the curve pairs determines thetransformation T. Below is a description of how this kind oftransformation functions in multi-component analysis of Hb varietieswith a plurality of different wavelengths and a plurality of differenthaemoglobin components.

The description is limited to the use of four different wavelengths.However, under certain assumptions, the procedure can be used to analysemore than four Hb varieties with sufficient accuracy.

In the analysing procedure of the invention, the total extinctioncoefficient of dyshaemoglobins at wavelength i is written as: ##EQU6##where

    DysHb=MetHb+HbCO+HbX                                       (4)

and HbX is a third dyshaemoglobin variety. In its general representationHbX can also be interpreted as an artificial dye concentration in blood.

With four different wavelengths, the unknown system of equations can bewritten in the form: ##EQU7## where the ε-matrix is in accordance withthe Lambert-Beer theory and contains the extinction coefficientsdocumented for blood in literature. The modulation percentages predictedby the theory are thus in the vertical %mod i vector. C again representsthe transformation of the units. The concentrations of different Hbvarieties are presented as proportional shares.

The experimentally measured modulation percentages are: ##EQU8##

By comparing equations 5 and 6, one observes that ##EQU9##

It can be observed at this point that the transformation T is neitherlinear nor e.g. a matrix multiplication, but it does have an unambiguousinverse transformation T⁻¹. We shall now calculate a representation ofthis transformation.

Equation (5) is divided into modulation ratios as follows: ##EQU10##where k,l=1, 2, 3, 4 and k #1. In the case of four wavelengths, thisdivision can be made in six different ways, i.e. there will be sixexpressions (8).

Let us write [(%mod k)/(%mod l)]=Z and solve each equation (8) for HbO2:##EQU11##

The theoretical modulation ratio Z can be expressed using anexperimentally measured modulation ratio Z' so that

    Z=f.sub.kl (Z')                                            (10)

which determines the transformation inv (T): Z'→Z. The experimentallycorrect fractional oxygen saturation is obtained from equation (9) bysubstituting formula (10) for Z. Consequently, when the wavelengths are660 nm and 940 nm, function f (660,940) transfers the experimentallymeasured curve in FIG. 7 onto the Lambert-Beer curve by changing thenumeric values on the vertical axis in the manner determined by functionf(660,940). Function f depends on the total absorption and scatteringappearing at wavelengths k and l, but no longer on different Hbvarieties. Thus, function f can be found e.g. via hypoxemia tests withnormal HbCO and MetHb concentrations. This makes the calibration offractional saturation measurement considerably easier. Function f isknown via the calibration data for conventional pulse oximetermeasurement at wavelengths R=660 nm and IR=900 nm (or 940 nm). If therepresentations of transformation inv(T) are designated e.g. atwavelengths 630 nm, 690 nm and 900 nm as follows: ##EQU12## then allfunctions f can be determined by changing the concentration of one Hbvariety and therefore the total absorption for the wavelength inquestion.

Representations R'-R, Q'-Q, P'-P, S'-S, T'-T and U'-U being known,equations (3) and (4) are solved and the fractional saturation level issolved for 3-6 modulation ratios from the system of equations (9)iteratively so that the fractional oxygen saturation level HbO2, totalconcentration of dyshaemoglobin DysHb and the composition ofdyshaemoglobin varieties can be determined in a compatible manner. Thecalculation may include verification of the result by checking itagainst the condition built into the equations that the sum of all Hbvarieties is 100%. The iteration process can be accelerated usingclinical information about the patient's condition. Typically, only oneDysHb variety in the patient's blood has an increased value. Moreover,the amount of dyshaemoglobin changes slowly with time, so equations (3)and (4) need not necessarily be used in real-time calculation of thefractional oxygen saturation level. These equations thus have a greaterimportance in the identification of dyshaemoglobin varieties.

Above, the way in which multi-component analysis of haemoglobincomposition of blood can be implemented has been described. Next, weshall discuss the question of how the concentration of dyshaemoglobincan be identified. After that there follows a discussion of how thesensor calibration data are saved.

In clinical measurement of fractional oxygen saturation, thedyshaemoglobin variety to be monitored is generally known beforehand. Itis generally also known that the concentrations of other dyshaemoglobinvarieties are normal and remain unchanged in the planned treatment. Insuch a situation, the user may input the dyshaemoglobin variety ofprimary interest into the apparatus. In this case, the non-interestingdyshaemoglobin concentrations are set as normal parameters inexpressions (3) and (4) while the dyshaemoglobin variety of primaryinterest is set as an unknown parameter. In principle, the user may alsomake haemoglobin values measured from a blood sample available to theapparatus. The apparatus performs the calculation using oxyhaemoglobin(HbO2), deoxyhaemoglobin (Hb) and the dyshaemoglobin variety of primaryinterest as unknown variables while the amount of other dyshaemoglobinvarieties is taken to be constant.

For reasons of user friendliness, it is generally required that itshould be possible to "measure" the dyshaemoglobin variety in the sameway as the amount of oxyhaemoglobin. In this case, the dyshaemoglobincomposition is not known in advance, and the effects of treatmentadministered to the patient are not known, either. In principle, thenormal composition of human blood, which contains four haemoglobinvarieties, can be analysed with four different wavelengths. In the bloodof a person who is critically ill, other HbX varieties or other dyes mayexist in consequence of treatment or medication. The existence of thesemay be identified by stipulating that the HbO2 value calculated viaiteration from six different modulation ratios should fall withincertain predetermined limits of variation. An additional stipulationthat may be applied is the assumption included in the equations that thesum of haemoglobin varieties should be 100%. Identification of adivergent HbX variety or dye is effected when the iteration result isnot consistent. If the result of iterative calculation indicates anincreased MetHb value and the patient is having NO-treatment, the HbXvariety is assumed to be nitrosylhaemoglobin HbNO. The extinctioncoefficient of this is added to expressions (3) and (4). The HbCO in theexpressions is set to the normal level or about 1%. If the nextiterative calculation yields a substantially better consistency, HbNO isidentified. If in normal iterative calculation no increased MetHb orHbCO values are observed but the accuracy of iteration is low, then acheck is made to verify whether any other pre-programmed Hb varieties ordyes exist. If the iterative calculation yields a substantially betterresult, then the haemoglobin variety or dye in question is identified.

Another advantage provided by the calculation method described above isthat changes in the sensor LED wavelengths can be easily and simplytaken into account in the calculation of fractional saturation. Thesensor calibration process comprises the following steps:

1. The wavelengths and spectral emission of the sensor LEDs aremeasured.

2. For the absorption curves of Hb varieties documented in theLambert-Beer theory or in literature, an effective extinctioncoefficient is calculated for the wavelength of each LED, taking intoaccount the medium wavelength and spectral line width of the LED. Thereceiver sensitivity curve can also be included in this process.

3. The extinction coefficients ε_(ij) are stored separately e.g. inmatrix format for all light elements.

4. The matrix of the extinction coefficients is stored in a storagedevice, such as an EEPROM, provided in the sensor.

5. Transformations of the theoretical extinction matrix and of anempirical matrix are stored in the same EEPROM as functions R, Q, P, S,T, and U. These functions are used in the procedure in the mannerdescribed above.

6. The sensor type, i.e. optionally MetHb sensor, CO sensor, etc., isstored in the EEPROM.

7. Data relating to the manufacture and guarantee of the sensor andother corresponding data are stored in the EEPROM.

Therefore, the sensor of the present invention provides a greatadvantage because the measuring apparatus can read the contents of theEEPROM and use the calibration data in the calculation of fractionalsaturation.

The invention is not limited to the examples of its embodimentsdescribed above, but instead many variations are possible within theframework of the inventive idea defined by the claims.

We claim:
 1. A method for non-invasively determining the amount of alight absorbing substance in the blood of a subject, the blood having atleast two light absorbing substances, said method comprising the stepsof:(A) applying light of a first wavelength to tissue of the subjectcontaining blood to obtain a first transmitted light quantity, the firsttransmitted light quantity having a pulsatile component and anon-pulsatile component; (B) applying light of a second wavelength totissue of the subject containing blood to obtain a second transmittedlight quantity, the second transmitted light quantity having a pulsatilecomponent and a non-pulsatile component; (C) ratioing the pulsatilecomponent of the first transmitted light quantity to the total firsttransmitted light quantity and ratioing the pulsatile component of thesecond light quantity to the total second transmitted light quantity toobtain first and second percentage modulation quantities (% mod 1, 2 . .. ); (D) ratioing the first and second percentage modulation quantities(% mod 1, 2 . . . ) to obtain a modulation ratio quantity (Z'); (E)formulating substance concentration determining equations (Eq 8, 9) forwavelengths corresponding to the first and second light wavelengths; (F)altering one of a modulation ratio term (Z', Z) or extinctioncoefficient terms (ε', ε) in the equations in accordance with thepercentage modulation quantities to account for the application of thelight of the first and second wavelengths to the tissue as well as theblood of the subject; (G) solving the equations for at least oneselected substance present in the blood; and (H) determining, from thesolution of the equations, the amount of the selected substance in theblood.
 2. The method according to claim 1, wherein step (F) is furtherdefined as applying a function (f) to the modulation ratio quantity (Z')obtained from the transmitted light modulation percentage quantities toobtain a modulation term (Z) for use in the solution of equationsformulated according to the Lambert-Beer law.
 3. The method according toclaim 1, wherein step (F) is further defined as applying an extinctioncoefficient transform (T) to substance extinction coefficients (ε_(ij))obtained in accordance with the Lambert-Beer law to provide haemoglobinextinction coefficients (ε'_(ij)) for blood in tissues and wherein step(G) is further defined as solving the equations using the modulationratio quantity term (Z') obtained from the transmitted light modulationpercentage quantities.
 4. The method according to claim 1, furtherdefined as a method for determining the relative amount of a given formof haemoglobin in the blood of the subject, the blood having at leasttwo forms of haemoglobin, wherein step (G) is further defined as solvingthe equations for at least one selected form of haemoglobin in theblood, and wherein step (H) is further defined as determining therelative amount of the given form of haemoglobin in the blood.
 5. Themethod according to claim 4, wherein step (G) is further defined assolving the equations for all haemoglobin forms present in the blood. 6.The method according to claim 4, further including the step ofiteratively solving the equations to determine the relative amount of aselected given form of haemoglobin.
 7. The method according to claim 4,further including the steps of iteratively solving the equations and asidentifying a selected given form of haemoglobin from the results of theiterative solutions.
 8. The method according to claim 4, further definedas iteratively solving the equations for a given form of haemaglobinpresent in the blood while formulating the equations in accordance withthe assumption that the amounts of one or more other forms ofhaemoglobin in the blood remain constant.
 9. The method according toclaim 4, further including the step of iteratively solving the equationsfor a given form of haemoglobin to determine changes over time in theamount of the given form of haemoglobin.
 10. The method according toclaim 1 wherein the method includes the steps of:applying light of fourdifferent wavelengths to the tissue of the subject to provide fourtransmitted light quantities; step (C) is further defined as ratioingthe pulsatile and total transmitted light quantities of each of thewavelengths to obtain four percentage modulation quantities; step (D) isfurther defined as ratioing the four percentage modulation quantities toobtain a plurality of modulation ratio quantities; step (E) is furtherdefined as formulating equations for the four different wavelengths; andstep (F) is further defined as altering the modulation ratio term (Z',Z) or extinction coefficient terms (ε', ε) to account for theapplication of light of the four wavelengths to the tissue, as well asthe blood of the subject.
 11. The method according to claim 10, furtherincluding the step of iteratively solving the equations and wherein oneor more modulation ratios derived from selected wavelengths is/areweighted in the iterative solution of the equations.
 12. The methodaccording to claim 10, further defined as a method for determining therelative amount of a given form of haemoglobin in the blood of thesubject, the blood having at least two forms of haemoglobin, whereinstep (G) is further defined as solving the equation for at least oneselected form of haemoglobin in the blood, and wherein step (H) isfurther defined as determining the relative amount of the given form ofhaemoglobin in the blood.
 13. The method according to claim 4, or claim12, further defined as including the step of selecting one of saidwavelengths so that there is a large difference between the extinctioncoefficients of selected forms of haemoglobin at the one wavelength andselecting another wavelength to be at or near the isobestic point of theextinction coefficients of the selected forms of haemoglobin.
 14. Themethod according to claim 4, or claim 12, further defined as includingthe step of selecting one of said wavelengths so that there is a largedifference of one sign between the extinction coefficients of selectedforms of haemoglobin at the one wavelength and selecting anotherwavelength so that there is a large difference of the opposite signbetween the extinction coefficients of the selected forms of haemoglobinat the another wavelength.
 15. The method according to claim 4, or claim12, wherein the wavelengths are selected in a range of 600 nm to 1000nm.
 16. The method according to claim 15, wherein the wavelengths areselected from the following ranges: 620-650; 655-665; 680-750; and790-1000 nm.
 17. The method according to claim 16, wherein thewavelengths are selected as follows: 645; 658; 690; and 900 nm.
 18. Amethod according to claim 16, wherein the wavelengths are selected asfollows: 632, 658, 690, and 900 nm.
 19. A method according to claim 15,wherein at least one of the wavelengths is below 660 nm.
 20. A methodaccording to claim 15, further including the step of iteratively solvingthe equations and wherein modulation ratio quantities derived fromwavelengths of 660 nm, 690 nm, and 900 nm are weighted with respect tomodulation ratio quantities derived from wavelengths in a range of620-650 nm.
 21. A method according to claim 15, further defined asdetermining the relative amount of at least one of methaemoglobin andcarboxyhaemoglobin in the blood and as using light having a wavelengthin a range of 620-650 nm.
 22. A method according to claim 4, or 12further defined as determining the relative amount of oxyhaemoglobin ordeoxyhaemoglobin in the blood.
 23. A method according to claim 4, or 12further defined as a method for determining the relative amount of aplurality of forms of haemoglobin and as including the step of applyinglight having a plurality of different wavelengths, the number ofdifferent wavelengths corresponding to the number of forms ofhaemoglobin, the relative amounts of which are to be determined.
 24. Amethod according to claim 23, further defined as a method fordetermining the relative amount of one or more of oxyhaemoglobin,deoxyhaemoglobin, and dyshaemoglobins in the blood.
 25. The methodaccording to claim 10, wherein the wavelengths are selected to optimizethe determination of a given form of haemoglobin.
 26. A method accordingto claim 1, wherein steps (A) and (B) are further defined as applyinglight to substantially the same blood containing tissue in both steps.27. A method according to claim 1, wherein the modulation ratio quantityand extinction coefficient terms in the equation are altered responsiveto small spectral changes in the wavelengths of the light.
 28. A methodaccording to claim 1, further defined as carrying out a calibrationprocedure and including the steps of obtaining modulation ratioquantities (Z') from the blood of the subject for the first and secondwavelengths; obtaining extinction coefficients and modulation ratioquantities (Z) under Lambert-Beer law conditions for the first andsecond wavelengths for blood having the same properties as the subject'sblood; comparing the modulation ratio quantities (Z, Z'); and using datafrom the comparison to calibrate the determination method.
 29. A methodaccording to claim 1, wherein steps (A) and (B) are further defined aspassing light through tissue of the subject.
 30. A method according toclaim 1, wherein steps (A) and (B) are further defined as reflectinglight off the tissue of the subject.
 31. A sensor for use in determiningthe amount of one or more light absorbing substances in the blood of asubject, said sensor comprising:a light source emitting light forapplication to blood containing tissue of the subject, the emitted lightof said source having at least two predetermined center wavelengths; areceiver receiving light from the tissue; a data storage device storingextinction coefficient data, the stored data including extinctioncoefficients separately determined for each predetermined centerwavelength and substance; and output means for providing output signalsfrom said sensor relating to characteristics of the light received bysaid receiver and the extinction coefficients stored in said datastorage device.
 32. A sensor according to claim 31, wherein said storagedevice is further defined as storing functions which are applied tosignal quantities resulting from the light received by said receiver toobtain output signal quantities suitable for use in the solution ofequations formulated according to the Lambert-Beer law.
 33. A sensoraccording to claim 31, wherein said data storage device is furtherdefined as storing an extinction coefficient transform appliable toextinction coefficients obtained in accordance with the Lambert-Beer lawto provide substance extinction coefficients for blood in tissue.
 34. Asensor according to in claim 31, wherein said storage device is furtherdefined as storing sensor identification information.
 35. A sensoraccording to claim 31, further defined as a sensor for determining therelative amount of one or more forms of haemoglobin in the blood of thesubject.
 36. A sensor according to claim 35, wherein said light sourcecomprises a plurality of light emitting elements emitting light havingpredetermined center wavelengths, at least one of which is below 660 nm,said light emitting elements emitting light in the following wavelengthranges; 620-650 nm, 655-665 nm, 680-750 nm, and 790-1000 nm.
 37. Asensor according to claim 31 further including filter means in the lightpath between said light source and said receiver for providing lighthaving said at least two predetermined center wavelengths.
 38. A sensoraccording to claim 35, including filter means in the light path betweensaid light source and said light receiver and wherein said filter meansprovides light of a plurality of predetermined center wavelengths, atleast one of which is below 660 nm, said filter means providing light inthe following wavelength ranges; 620-650 mn, 655-665 nm, 680-750 nm, and790-1000 nm.
 39. A sensor according to claim 31, wherein said lightsource comprises a first set of light emitting elements and a second setof light emitting elements, and wherein said first and second sets oflight emitting elements emit light in one common wavelength range.
 40. Asensor according to claim 31, wherein said light source and saidreceiver are arranged so that said receiver receives light transmittedthrough the blood containing tissue.
 41. A sensor according to claim 31,wherein said light source and said receiver are arranged so that saidreceiver receives light reflected from the blood containing tissue. 42.A sensor according to claim 31, wherein said light source produces heatand is mounted on a heat sink for maintaining the temperature of saidlight source below a predetermined limit.
 43. A sensor according toclaim 31, wherein said light source produces heat, wherein said sensorhas a frame, and wherein said light source is fastened to said frame forconducting heat from said light source to said frame to maintain thetemperature of the light source below a predetermined limit.
 44. Asensor according to claim 31, wherein said light source has a mountingframe and wherein said light source is located in a cavity of saidframe.
 45. A sensor according to claim 31, further including at leastone light transmitting, optical fiber coupled to at least one of saidlight source or receiver.
 46. A sensor according to claim 31, furtherincluding attachment means for attaching the sensor to the body of thesubject.
 47. Measuring apparatus for determining the amount of one ormore light absorbing substances in the blood of a subject, saidapparatus comprising:a light source emitting light for application toblood containing tissue of the subject, the emitted light of said sourcehaving at least two predetermined center wavelengths; a receiverreceiving light from the tissue and providing output signals responsiveto the received light; a data storage device storing extinctioncoefficient data, the stored data including extinction coefficientsseparately determined for each predetermined center wavelength and lightabsorbing substance; means coupled to said receiver for processing theoutput signals from the receiver; reader means coupled to said datastorage device for obtaining extinction coefficient data; andcalculating means coupled to said processing means and said reader meansfor determining the amount of light absorbing substance in the blood.48. Measuring apparatus according to claim 47, further defined as onefor determining the relative amount of one or more forms of haemoglobinin the blood of a subject.
 49. Measuring apparatus according to claim48, further defined as one for determining the relative amount of aplurality of forms of haemoglobin in the blood of the subject andwherein said calculating means is further defined as carrying out thedetermination of the forms of haemoglobin on a time division basis. 50.Measuring apparatus according to claim 47, wherein said data storagedevice stores sensor identification information for provision to saidreader means.